Implantable stimulation devices are devices that generate and deliver electrical stimuli to nerves and tissues for the therapy of various biological disorders, such as pacemakers to treat cardiac arrhythmia, defibrillators to treat cardiac fibrillation, cochlear stimulators to treat deafness, retinal stimulators to treat blindness, muscle stimulators to produce coordinated limb movement, spinal cord stimulators to treat chronic pain, cortical and deep brain stimulators to treat motor and psychological disorders, and other neural stimulators to treat urinary incontinence, sleep apnea, shoulder subluxation, etc. The description that follows will generally focus on the use of the invention within a Spinal Cord Stimulation (SCS) system, such as that disclosed in U.S. Pat. No. 6,516,227. However, the present invention may find applicability in any implantable medical device system.
As shown in FIGS. 1A and 1B, a SCS system typically includes an Implantable Pulse Generator (IPG) 10, which includes a biocompatible device case 12 formed of metallic material such as titanium for example. The case 12 typically houses the circuitry and battery 14 (FIG. 2B) necessary for the IPG to function, although IPGs can also be powered via external RF energy and without a battery. The IPG 10 is coupled to electrodes 16 via one or more electrode leads (two such leads 18 are shown), such that the electrodes 16 form an electrode array 20. The electrodes 16 are carried on a flexible body 22, which also houses the individual signal wires 24 coupled to each electrode. In the illustrated embodiment, there are eight electrodes on each lead, although the number of leads and electrodes is application specific and therefore can vary. The leads 18 couple to the IPG 10 using lead connectors 26, which are fixed in a header 28 comprising epoxy for example, which header is affixed to the case 12. In a SCS application, distal ends of electrode leads 18 are typically implanted on the right and left side of the dura within the patient's spinal cord. The proximal ends of leads 18 are then tunneled through the patient's tissue 100 to a distant location such as the buttocks where the IPG 10 is implanted, where the proximal leads ends are then connected to the lead connectors 26.
As shown in cross section in FIG. 2B, the IPG 10 typically includes an electronic substrate assembly including a printed circuit board (PCB) 30 containing various electronic components 32 necessary for operation of the IPG 10, some of which are described subsequently. Two coils are generally present in the IPG 10: a telemetry coil 34 used to transmit/receive data to/from an external controller (not shown); and a charging coil 36 for charging or recharging the IPG's battery 14 using an external charger 70 (FIG. 2A). These coils 34 and 36 are also shown in the perspective view of FIG. 1B, which omits the IPG case 12 for easier viewing. Although shown as inside in the case 12 in the Figures, the telemetry coil 34 can alternatively be fixed in header 28. Coils 34 and 36 may alternative be combined into a single telemetry/charging coil.
FIG. 2A shows a plan view of the external charger 70, and FIG. 2B shows it in cross section and in relation to the IPG 10 during a charging session. The external charger 70 is used to wirelessly provide operational power to the implantable medical device, such as to charge or recharge the IPG's battery 14, and includes at least one PCB 72 (two are shown; see U.S. Patent Application Publication 2008/0027500); electronic components 74, some of which are subsequently discussed; a charging coil 76; and a battery 78 for providing operational power for the external charger 70 and for the production of a magnetic charging field 80 from the coil 76. These components are typically housed within a case 77 which is sized to be hand held and portable, which may be made of plastic for example.
The external charger 70 has a user interface 82, which typically comprises an on/off switch 84 to activate the production of the magnetic charging field 80; an LED 86 to indicate the status of the on/off switch 84; and a speaker 88. The speaker 88 emits a “beep” for example if the external charger 70 detects that its charging coil 76 is not in good alignment with the charging coil 36 in the IPG 10 during a charging session, as discussed further below. The external charger 70 may be placed in a pouch around a patient's waist to position the external charger 70 in alignment with the IPG 10 during a charging session. Typically, the external charger 70 is touching the patient's tissue 100 during a charging session as shown, although the patient's clothing or the material of the pouch may intervene.
Wireless power transfer from the external charger 70 to the IPG 10 occurs by magnetic inductive coupling between coils 76 and 36. Referring to FIG. 3, when the external charger 70 is activated (e.g., on/off switch 84 is pressed), a charging circuit 94 under control of control circuitry 92 (e.g., a microcontroller) energizes coil 76 with a non-data-modulated AC current (Icharge) to create a magnetic charging field 80. The frequency of the magnetic charging field 80 may be on the order of 80 kHz for example, and may be set by the inductance of the coil 76 and the capacitance of a tuning capacitor (not shown). The magnetic charging field 80 induces a current in the IPG 10's charging coil 36, which charging circuitry in the IPG 10 uses to provide operational power to the IPG 10. For example, in an IPG with a rechargeable battery 14, charging circuitry can include a rectifier 44 to produce a DC voltage used to provide a charging current (Ibat) to recharge the IPG's battery 14, perhaps using an intermediate charge control and battery protection circuit 46 as shown. Wireless power transfer occurs transcutaneously through the patient's tissue 100.
The IPG 10 can also communicate data back to the external charger 70 along link 81 using Load Shift Keying (LSK) telemetry. Relevant data, such as the capacity of the battery, is sent from control circuitry 38 in the IPG 10 (e.g., a microcontroller) to a LSK modulator 40, which creates a series of digital data bits (LSK data 48). This data is input to the gate of a load transistor 42 to modulate the impedance of the charging coil 36 in the IPG 10. Such modulation of the charging coil 36 is detectable at the external charger 70 due to the mutual inductance between the coils 76 and 36, and will change the magnitude of the AC voltage needed at coil 76 (Vcoil) to drive the charging current, Icharge. If coil 36 is shorted (LSK data=1), Vcoil increases (Vcoil1) to maintain Icharge; if not shorted (LSK data=0), Vcoil decreases (Vcoil0), as shown in the waveform in FIG. 5. LSK demodulator 96 (receiver circuitry) in the external charger 70 can detect these changes in Vcoil (ΔV) to recover the series of digital data bits, which data is then received at control circuitry 92 so that appropriate action can be taken, such as ceasing production of the magnetic charging field 80 (i.e., setting Icharge to zero) when the battery 14 in the IPG 10 is fully charged. Note that the nature of LSK telemetry as described here only allows for telemetry from the IPG 10 to the external charger 70 when a magnetic charging field 80 is being produced, and requires the external charger 70 to be close the IPG 10. See, e.g., U.S. Patent Application Publication 2013/0123881 for further details regarding the use of LSK telemetry in an external charger system.
It is generally desirable to charge the IPG's battery 14 as quickly as possible to minimize inconvenience to the patient. One way to decrease charging time is to increase the strength of the magnetic charging field 80 by increasing Icharge in the charging coil 76 of the external charger 70. Increasing the magnetic charging field 80 will increase the current/voltage induced in the coil 36 of the IPG 10, which increases the battery charging current, Ibat, hence charging the battery 14 faster.
However, the strength of the magnetic charging field 80 can only be increased so far before heating becomes a concern. Heating is an inevitable side effect of inductive charging using magnetic fields, and can result because of activation of relevant charging circuitry in the external charger 70 or IPG 10, or as a result of eddy currents formed by the magnetic charging field 80 in conductive structures in either device. Heating is a safety concern. The external charger 70 is usually in contact with the patient's tissue 100 during a charging session, and of course the IPG 10 is inside the patient. If the temperature of either exceeds a given safe temperature, the patient's tissue may be aggravated or damaged.
The alignment between the external charger 70 and the IPG 10 can affect heating, as shown in FIGS. 4A and 4B. In FIG. 4A, the charging coils 76 and 36 in the external charger 70 and the IPG 10 are well aligned, because the axes 76′ and 36′ around which the coils 76 and 36 are wound are collinear. As such, these coils 76 and 36 are well coupled electrically, meaning that a higher percentage of the power expended at coil 76 in creating the magnetic charging field 80 is actually received at coil 36, which leads to higher values for Ibat. In FIG. 4B, the charging coils 76 and 36 are laterally (radially) misaligned (r), which reduces the electrical coupling between the coils. Increasing the vertical distance d between the coils 76 and 36 (FIG. 4C), or increasing the angle (A) between the preferably parallel planes in which they reside (FIG. 4D), will also reduce coupling.
If it is desired that the alignment scenarios of FIGS. 4A and 4B charge the battery 14 at the same rate (Ibat=Y), then a higher value for Icharge (Icharge>X) will be needed in the misaligned scenario of FIG. 6B compared to the well-aligned scenario of FIG. 4A (Icharge=X). A higher value for Icharge in FIG. 4B will create a more intense magnetic charging field 80 that tends to increase the temperature of the environment (T>Z) when compared to the temperature of the environment in FIG. 4A (T=Z). If it is desired that the temperature be the same for both scenarios, then Icharge can be lowered in FIG. 4B, but this will also lower Ibat, and hence the battery 14 in that scenario would take longer to charge. In short, misalignment between the external charger 70 and the IPG 10 is not desired.
Accordingly, the art has disclosed several manners for determining misalignment between an external charger and an IPG, which techniques usually result in some form of user-discernible output letting the patient know when alignment is poor (such as via speaker 88 discussed earlier). Such techniques may also inform a patient how to fix the alignment, such as by indicating a direction the external charger should be moved relative to the IPG 10. See, e.g., U.S. Pat. Nos. 8,473,066 and 8,311,638.
Previous external charger alignment techniques however are difficult to implement, and may not precisely determine alignment as they rely on inferences gleaned from electrical measurements taken at the external charger during the charging session. For example, one prior art alignment techniques relies on determining the loading of the charging coil in the external charger during production of the magnetic charging field. Specifically, the voltage across the charging coil (Vcoil) is reviewed at the external charger and compared to a Vcoil threshold to determine alignment. This technique though suffers in its inability to distinguish between the scenarios of FIGS. 4B and 4C for example. In either of these scenarios, Vcoil would be higher due to poor coupling, but in FIG. 4B the poor coupling arises from misalignment (r), whereas in FIG. 4C the alignment is as good as it can be given the IPG 10's depth (d). A modification to this technique helpful in distinguishing these scenarios requires transmitting the magnetic charging field at different frequencies and measuring the input current to the charging coil in the external charger to estimate an implant depth (d), and thus to set an appropriate Vcoil threshold. See, e.g., U.S. Patent Application Publication 2010/0137948. However, the additional overhead of having to produce magnetic charging fields at different frequencies makes this technique complicated.
Other alignment techniques require the external charger to have positioning coils in addition to the main charging coil (e.g., 76), which positioning coils are used to sense magnetic fields in the environment. In these techniques, measurements taken from the positioning coils during the charging session are used to determine misalignment, and to indicate a direction the external charger can be moved to improve alignment (coupling). See, e.g., U.S. Pat. Nos. 8,473,066 and 8,311,638. But these positioning-coil measurements again rely on loading, and therefore are indirect. Moreover, assessing the loading of the position coils does not necessarily discriminate between loading caused by coupling of the charging coil in the IPG, and coupling caused by other sources, such as the conductive material used for the IPG's case. Moreover, while positioning coils can provide a general sense of the misalignment direction between the external charger and the IPG, they do not compute a misalignment distance—i.e., how far the external charger must be moved in the misalignment direction to achieve good alignment. Such data would be useful to the user who is attempting to improve external charger alignment with her IPG. A more accurate means of determining external charger/IPG alignment is therefore desired.